Image-forming MR methods which utilize the interaction between magnetic fields and nuclear spins in order to form two-dimensional or three-dimensional images are widely used nowadays, notably in the field of medical diagnostics, because for the imaging of soft tissue they are superior to other imaging methods in many respects, do not require ionizing radiation and are usually not invasive.
According to the MR method in general, the body of the patient to be examined is arranged in a strong, uniform magnetic field B0 whose direction at the same time defines an axis (normally the z-axis) of the co-ordinate system to which the measurement is related. The magnetic field B0 produces different energy levels for the individual nuclear spins in dependence on the magnetic field strength which can be excited (spin resonance) by application of an electromagnetic alternating field (RF field) of defined frequency (so-called Larmor frequency, or MR frequency). From a macroscopic point of view the distribution of the individual nuclear spins produces an overall magnetization which can be deflected out of the state of equilibrium by application of an electromagnetic pulse of appropriate frequency (RF pulse) while the corresponding magnetic field B1 of this RF pulse extends perpendicular to the z-axis, so that the magnetization performs a precessional motion about the z-axis. The precessional motion describes a surface of a cone whose angle of aperture is referred to as flip angle. The magnitude of the flip angle is dependent on the strength and the duration of the applied electromagnetic pulse. In the case of a so-called 90° pulse, the magnetization is deflected from the z axis to the transverse plane (flip angle 90°).
After termination of the RF pulse, the magnetization relaxes back to the original state of equilibrium, in which the magnetization in the z direction is built up again with a first time constant T1 (spin lattice or longitudinal relaxation time), and the magnetization in the direction perpendicular to the z direction relaxes with a second and shorter time constant T2 (spin-spin or transverse relaxation time). The transverse magnetization and its variation can be detected by means of receiving RF coils which are arranged and oriented within an examination volume of the MR device in such a manner that the variation of the magnetization is measured in the direction perpendicular to the z-axis. The decay of the transverse magnetization is accompanied by dephasing taking place after RF excitation caused by local magnetic field inhomogeneities facilitating a transition from an ordered state with the same signal phase to a state in which all phase angles are uniformly distributed. The dephasing can be compensated by means of a refocusing RF pulse (for example a 180° pulse). This produces an echo signal (spin echo) in the receiving coils.
In order to realize spatial resolution in the body, constant magnetic field gradients extending along the three main axes are superposed on the uniform magnetic field B0, leading to a linear spatial dependency of the spin resonance frequency. The signal picked up in the receiving coils then contains components of different frequencies which can be associated with different locations in the body. The signal data obtained via the receiving coils correspond to the spatial frequency domain and are called k-space data. The k-space data usually include multiple lines acquired of different phase encoding. Each line is digitized by collecting a number of samples. A set of k-space data is converted to an MR image by means of Fourier transformation.
It is important to note that transverse magnetization dephases also in presence of constant magnetic field gradients. This process can be reversed, similar to the formation of RF induced echoes, by appropriate gradient reversal forming a so-called gradient echo. However, in case of a gradient echo, effects of main field inhomogeneities, chemical shift and other off-resonances effects are not refocused, in contrast to the RF refocused echo.
In MR imaging, it is often desired to obtain information about the relative contribution of different chemical species, such as water and fat, to the overall signal, either to suppress the contribution of some of them or to separately or jointly analyze the contribution of all of them. These contributions can be calculated if information from two or more corresponding gradient echoes, acquired at different echo times, is combined. This may be considered as chemical shift encoding, in which an additional dimension, the chemical shift dimension, is defined and encoded by acquiring a couple of images at slightly different echo times. In particular for water/fat separation, these types of experiments are often referred to as Dixon-type of measurements. By means of Dixon imaging or Dixon water/fat imaging, a water/fat separation can be achieved by calculating contributions of water and fat from two or more corresponding echoes, acquired at different echo times. In general such a separation is possible because there is a known precessional frequency difference between hydrogens in fat and water. In its simplest form, water and fat can be seen as forming a 2-line spectrum (which does not exclude more sophisticated spectral models). Thus, water and fat images are generated by either addition or subtraction of the ‘in phase’ and ‘out of phase’ datasets, obtained at defined gradient echo times.
It is known that a number of artifacts can appear in MR images acquired from a region of the examined body in which flow phenomena occur, for example an arterial blood flow.